1. Field of the Invention
This invention generally relates to a positron emission tomography (PET) imaging system, more specifically, to pairing and filtering detection events for energy, timing, and spatial characteristics.
2. Discussion of the Background
The use of positron emission tomography (PET) is growing in the field of medical imaging. In PET imaging, a radiopharmaceutical agent is introduced into the object to be imaged via injection, inhalation, or ingestion. After administration of the radiopharmaceutical, the physical and bio-molecular properties of the agent will cause it to concentrate at specific locations in the human body. The actual spatial distribution of the agent, the intensity of the region of accumulation of the agent, and the kinetics of the process from administration to eventually elimination are all factors that may have clinical significance. During this process, a positron emitter attached to the radiopharmaceutical agent will emit positrons according to the physical properties of the isotope, such as half-life, branching ratio, etc.
The radionuclide emits positrons, and when an emitted positron collides with an electron, an annihilation event occurs, wherein the positron and electron are destroyed. Most of the time, an annihilation event produces two gamma rays at 511 keV traveling at substantially 180 degrees apart.
By detecting the two gamma rays, and drawing a line between their locations, i.e., the line-of-response (LOR), one can retrieve the likely location of the original disintegration. While this process will only identify a line of possible interaction, by accumulating a large number of those lines, and through a tomographic reconstruction process, the original distribution can be estimated. In addition to the location of the two scintillation events, if accurate timing (within few hundred picoseconds) is available, a time-of-flight (TOF) calculation can add more information regarding the likely position of the event along the line. Limitations in the timing resolution of the scanner will determine the accuracy of the positioning along this line. Limitations in the determination of the location of the original scintillation events will determine the ultimate spatial resolution of the scanner, while the specific characteristics of the isotope (e.g., energy of the positron) will also contribute (via positron range and co-linearity of the two gamma rays) to the determination of the spatial resolution the specific agent.
The collection of a large number of events creates the necessary information for an image of an object to be estimated through tomographic reconstruction. Two detected events occurring at substantially the same time at corresponding detector elements form a line-of-response that can be histogrammed according to their geometric attributes to define projections, or sinograms to be reconstructed. Events can also be added to the image individually.
The fundamental element of the data collection and image reconstruction is therefore the LOR, which is the line traversing the system-patient aperture. Additional information can be obtained regarding the location of the event. First, it is known that, through sampling and reconstruction, the ability of the system to reconstruct or position a point is not space-invariant across the field of view, but is better in the center, slowly degrading toward the periphery. A point-spread-function (PSF) is typically used to characterize this behavior. Tools have been developed to incorporate the PSF into the reconstruction process. Second, the time-of-flight, or time differential between the arrival of the gamma ray on each detector involved in the detection of the pair, can be used to determine where along the LOR the event is more likely to have occurred.
The above described detection process must be repeated for a large number of annihilation events. While each imaging case must be analyzed to determine how many counts (i.e., paired events) are required to support the imaging task, current practice dictates that a typical 100-cm long, FDG (fluoro-deoxyglucose) study will need to accumulate several hundred million counts. The time required to accumulate this number of counts is determined by the injected dose of the agent and the sensitivity and counting capacity of the scanner.
PET imaging systems use detectors positioned across from one another to detect the gamma rays emitting from the object. Typically a ring of detectors is used in order to detect gamma rays coming from each angle. Thus, a PET scanner is typically substantially cylindrical to be able to capture as much radiation as possible, which should be, by definition, isotropic. The use of partial rings and rotation of the detector to capture missing angles is also possible, but these approaches have severe consequences for the overall sensitivity of the scanner. In a cylindrical geometry, in which all gamma rays included in a plane have a chance to interact with the detector, an increase in the axial dimension has a very beneficial effect on the sensitivity or ability to capture the radiation. Thus, the best design is that of a sphere, in which all gamma rays have the opportunity to be detected. Of course, for application to humans, the spherical design would have to be very large and thus very expensive. Accordingly, a cylindrical geometry, with the axial extent of the detector being a variable, is realistically the starting point of the design of a modern PET scanner.
While a PET detector can only detect single interactions, i.e., one gamma ray interacting with a crystal and generating light through a scintillation process, PET events are defined by two of those detections occurring at substantially the same time or in coincidence, at substantially 511 keV, and in a geometry compatible with the annihilation event to have occurred in an object of interest. It is therefore required for a PET system to properly identify the timeline for each event in order to correctly match or pair events. This is typically accomplished by constructing a complex network of real-time comparators. As the requirement for count rate is also very demanding (up to hundreds of millions of single events per second), the construction of the coincidence circuitry also needs to handle a very large numbers of counts.
Because of the high demand on efficiency, i.e., being able to receive and process hundreds of millions of events per second, the design of the coincidence circuitry is typically one of the most important elements of the PET detection system. Trigger lines are typically brought to centralized hardware for comparison. Usually the coincidence window, or the period of time within which two events will be deemed to be “at the same time,” is set from high-level system controls and does not typically vary during a study or even between studies.
Conventional PET systems suffer from several disadvantages and limitations. For example, conventional systems are very complex since the number of possible coincidences that can originate from independent detectors increases exponentially. While this complexity is manageable when trigger signals come from a few dozens or even a few hundreds of detector elements, it can become simply intractable with pixelated systems that can count several thousand independent signals.
Further, conventional PET systems are also rigid and allow for very little variation in the bandwidth, geometry, and filtering parameters. In addition, coincidence circuitry in conventional systems is typically destructive in the sense that the very action of pairing events destroys the timing information, and a variable coincidence window cannot be applied on the same data.